Blood cell analyzer

ABSTRACT

A system for differentiating among white blood cells by flowing the latter in a supporting liquid through an elongated sheathedfluid flowcell in which various cell parameters are determined by illuminating the cells with radiation incident from at least three directions each being substantially more than 60* from one another, typically as a cone of radiation having a solid angle of much greater than 90*. Such illumination serves to minimize variations, due to orientation, in the subsequent radiation emitted by the cell.

U llited fStates'Patent {-191 Hirsch feld f 3,819,270 [451 time 25, 1974 BLOOD CELL ANALYZER [75] Inventor: Tomas Hirschfeld, Framingham,

Mass. v

[73] Assignee: Block Engineering, Inc., Cambridge,

, v Mass.

[22] Filed: Oct. 2, 1972 [21] Appl. No.: 294,245

[52] us. Cl. 356/39, 356/103 [51] Int. Cl. G0ln 33/16 [58] Field of Search....., 356/36, 39, 102, 103 [56] References Cited UNITED STATES PATENTS 3,315,229 4/1967 'Smithline 3356/36 3,662,176v 5/1972 Kamemsky eta'l. 356/39 3,669,542 6/1972 Capellaro 356/36 3,675,768 7/1972 Legorreta-Sanchez 356/39 3,710,933 l/1973 Fulwyler 356/39 OTHER PUBLICATIONS John Strong, Concepts of Classical Optics, Freeman & Co.,'San Francisco, 1958, pp. 353-355.

Primary ExaminerRonald L. Wibert Attorney, Agent, or Firm-Schiller & Pandiscio [57] ABSTRACT other, typically as a cone of radiation having a solid angle of much greater than 90. Such illumination serves to minimize variations, due to orientation, in

the subsequent radiation emitted by the cell.

7'Claims', 6 Drawing Figures PATENTEDJIIII25 I974 3.8 191270 sum 1 F 2 DIALYSIS DYE SOLUT|QN DI LUENT SAMPLE MIXING SHEATH SOURCE CHAMBER DIALYZER DILUTIOR FLU) TO DISCHARGE PUMP PUMP FIG. i \54 52 86? as 64 so 7 f(X) 84 1 f(X) I X 2 '6 8 TI F T 7; w- 64@ 9. o I; I: 5 a 1 9, 64

s E a REF REF REF 5 I: E 2' Z 1 a E 8 O. F (D 9 97 r 98 ET 8 I- I- COMP'\ COMP COMP l y 64 so COUNT I COUNT COUNT #l #2 #3 J (-IOO lOl QIOZ TO DISCHARGE termining the classification of blood cells.

The classification by a pathologist of a population of blood cells is usually based upon five visual parameters determined microscopically;-the color, size and shape of the cell nucleus after appropriate staining, and the color and sizeof stained cytoplasm. While the tech nique provides quite good classification, it is usually limited to a population of a few hundred'cells. Despite its high degree of redundancy in cell identification, the technique can nevertheless be in serious error, because the limited populations may not provide a statistically reliable sampling. 1

Considerable efforts have been made to provide apparatus that will automatically recognize the same parametersemployed by the pathologists. For example, a number of systems employ electronic image tubes'with computers ,to achieveip'attem recognition. However,- such'systems invariably require a very large number of resolution 'elernent swith a commensurately sized computer memory, andare therefore usually very expensive and often bulky and complex. I A number of automatic blood classification systems use non-imaging techniques in which the classification methods are based typically on size only or on color only. However, in such devices, the random orientation of cells leads'to severe. difficulties in measurement by radiometry, and thus tends to introduce serious inaccuracies in cell classification.

Because the transmission of light through optically thin particles illuminated by a beam wider than the maximum cross-section of the particle is independent of particle orientation, it has been a practice to limit photometry of stained cells to optically thin cases obtained by appropriate selection of dye, dye concentration, or wavelength. However, such an approach leads to serious signal-to-noise problems, as it requires measuring a small diminution of a bright background signal. In effect, such an approach trades one source of error (orientation dependence of transmission of optically thick particles) for another (measurement of a small signal in the presence of noise).

To overcome this latter problem, fluorescent dyes have been commonly used. If the dye isapplied such that the dyed particle is optically thin in the dyes absorption band and the particle is illuminated such that the detector only sees fluorescent emission from the .particle, both the background and the orientation problems are surmounted. Such techniques, however, give rise to a number of other problems. First, only a small percentage of dyes are fluorescennand of the more important dyes used in cell classification many are not fluorescent. Secondly, there appears to be little or no chemical or clinical information regarding the biological properties of the vast majority of fluorescent. dyes. To determine such properties and their diagnostic efficacy would require a research program of major magnitude. Such research must also concern itself with quenching of fluorescence and energy transfer as well. Finally, the nature of dyes is such thattheir absorption bands (and, in the case, of fluorescent dyes,.their"fl uo-,

rescent bands as well) are both broad (on the order of 500 A) and restricted to a limited region of the spectrum (on the order of 10,000 A'wide). If dyes are to be used in combination, their spectra must be well separated a condition which is vastly more difficult in the case of fluorescent dyes, since both absorption and fluorescent bands of any dye used in combination must not interfere with the absorption and fluorescent bands of any of the other dyes in that combination.

A principal object of the present invention is therefore to provide a novel system for classifying a number of blood cells. Another object of the invention is to provide a non-imaging measurement of cell parameters, compatible with standard hematological practice. Yet another object of the present invention is to provide such a non-imaging measurement of blood cell parameters wherein data is preprocessed so that the use of large computers is avoided.

Yet another objects of the present invention are to provide such a measurement system which can be employed with blood cells moving in a flow stream; to provide such a systemwherein measurements are made with minimized orientation-caused error; to provide such a system wherein a shape factor of a cell or cell nucleus is determined from simultaneous measurement of the magnitudes of two different functions of shape; to provide such a system wherein the shape factor is determined from simultaneous measurement of magnitudes related respectively to the volume of the nucleus and the surface area of the nucleus; and to provide such a system wherein the shape factor is determined from simultaneous measurement of magnitudes related respectively to the effective thickness and the volume of the cell nucleus. e i

- Yet another important object of the present invention is to provide a novel measurement system wherein measurements are made that minimizes variation caused by orientation whereby measurements of parameters of optically thick, dyed particles may be reproducibly made.

Yet other objects of the invention will in part be obvious and will in part appear hereinafter. The invention accordingly comprises the apparatus possessing the construction, combination of elements and arrangement of parts, and the method comprising the several steps and relation of one or more of such steps with re spect to each of the others, all of which are exemplified in the following detailed disclosure, and the scope of the application of which will be indicated in the claims.

For a fuller understanding of the nature and objects of the present invention, reference should be had to the following detailed description taken in connection with the accompanying drawings wherein like numerals denote like parts, and wherein;

FIG. 1 is a schematic block diagram illustrating application of the invention to a system for obtaining a differential count among selected types of blood cells;

FIG. 2 is a schematic diagram showing an exemplary optical system useful in part of the system'of FIG. 1;

FIG. 3 is a schematic diagram showing another exemplary optical system useful in connection with the system of FIG. 1',

FIG. 4 is a schematic diagram of an electrical circuit which is a modification of a portion of the system of FIG. 1 for alternative structure;

FIG. 5 .is a schematic diagram of a vibration on the optical system employed to minimize orientation prob lems; and

FIG. 6 is a schematic elevational diagram, partly in cross-section, showing a fragment of a preferred optical system for an orientation insensitive measuring system.

The system of the invention herein disclosed is com-- patible with standard hematology laboratory practice, and does not require cell fractionation or selective lysis and requires few solutions. The sample is not destroyed and is available for subsequent microscopic examination or other diagnostic tests. Most importantly, the present invention permits the direct assignment of cells to the major cell fractions. This reduces error, since no counts are derived as a remainder when two other counts are differenced, as for example, in a scheme which characterizes monocytes as being nongranulocytes which are also non-lymphocytes.

The term function" as used herein is intended to be interpreted in the mathematical sense to mean a variable, the value or magnitude of which is determined by a second variable. The term different functions is intended then to mean a plurality of functions wherein the laws of dependence on the second variables are different.

The present invention generally is a system comprising means for substantially simultaneously measuring the magnitudes of two different shape-dependent functions of the cell or its nucleus and means for determining a relationship between the magnitudes. For example, the property of sphericity can be considered to be the desired shape factor. For a perfect sphere, there are a number of different shape dependent functions, typically the surface area (A), the volume (V), the mean cross-sectional area (X) and the mean thickness (T). These latter can be defined respectively as A =7rd V =1rlP/6 X='rr(d/2) T 1r(2/3)d where d is the spherical diameter.

The relationship between any two of these shape dependent functions can be expressed by a ratio where n and m are exponents selected so that d, which is size dependent primarily, vanishes. Thus, for example, iff(d) is V, (d) is S, m= and n=%, the shape factor y for a cellular nucleus having a perfectly spherical shape will be a readily computed limit quite independent of the value of d. Any departures measured from such limit are measures of variation of the shape of the nucleus from sphericity and hence a determination of an aspect of shape. It will readily be seen that, for example, where f(d) is V and (Md) is T, by selecting m l and n /3, the shape factor y will again have a readily computcd limit representing a sphere. and will vary according to departures from sphericity.

In one embodiment of the present invention means are provided for deriving a magnitude which is proportional to the volume of a cell nucleus, preferably by measuring fluorescent re-emission of light absorbed by the nuclear DNA, and simultaneously deriving a magnitude related to the nuclear surface area, preferably by measuring the light scattering from the nucleus. Means are also provided for then determining a shape factor which is proportional to a ratio of the volumetrically related magnitude to the surface related magnitude.

In another embodiment of the invention, the two magnitudes derived are respectively related to other shape-dependent factors, particularly the apparent thickness and the volume of a cell nucleus, and a shape factor is obtained from a ratio of these two functions. Measurement here can be achieved from light absorption or by fluorescent re-emission. The effective thickness is measured by examining the self-shadowing reduction in the measured volume-at wavelengths where the particle is optically thick. Such reduction arises from the non-linearity of the transmission-thickness relationship.

Referring-now to the drawing, the system of FIG. 1 includes source 20 of a sample of blood cells. Source 20 is connectable through conduit 22 and valve 24 to a mixing chamber 26. The latter also is connectable through conduit 28 and valve 30 to supply 32 of a dye or stain bath. The output of chamber 26 is coupled to an input of dialyzer 34, another input of the latter being connectable through conduit 36 and valve 38 to a reservoir or supply 40 of dialysis solution. One output of dialyzer is for discharge of dialyzed stain, the other output of the dialyzer being connected as one input to dilution chamber 42. Another input to chamber 42 is connectable through conduit 44 and valve 46 to supply 48 of diluent fluid.

The diluent from supply 48 is preferably selected (and metered by valve 46) to provide a number of basic qualities for purposes of this invention. The diluent should, when added in metered proportion to the stained cell suspension from dialyzer 34, provide matching of the indices of refraction of the cell cytoplasm and the mixed fluids; because it is highly desirable to maintain substantially laminar flow through the flowcell at a high rate of flow, the diluent should also be selected so that, when added in metered proportion to the cell suspension, it will adjust the fluid viscosity to permit high speed laminar flow. The diluent also should be selected so that, when added to the cell suspension it will attain an optimum osmotic pressure with respect to the cells to maintain their stability. In some instances, pumps may be provided to yield the high rate of flow through the flow cell. In such case, the diluent may also serve to adjust osmotic pressure and thereby compensate for the high static pressure caused by the pumps.

The fluid output from dilution chamber 42 is connected to pump 50, and the output of the latter in turn is connected preferably to a central injector nozzle 51 of a sheathed-stream flow cell 52. The latter typically can be of the design disclosed in Advances in Automated Analysis, Technicon International Congress 1970, Volume l Clinical (1971) on pages 454-455 of the article by Alex M. Saunders et al, entitled A Rapid Automated System for Differentiating and Counting White Blood Cells." The annular space 53 around central injector 51 is connected to the output of second pump 54. The inputs to the latter is connected to a supply 56 of sheath fluid. Injector 51 and space 53 are disposed the fluid carrier becomes many times smaller than the dye concentration on a typical cell. The sample is then diluted in chamber 62 with diluent-from supply 48 to provide adequate separation between blood cells in flow cell 52. It is possible that thedilution alone may be adequate to reduce the solution concentration of dye, in which case the dialyzer may be eliminated. It is desirable to eliminate the dialyzer, asit accoiints for a major portion of, the totaltime required for measuring the parameters of a sample (the time between samples can, of course, be much smaller than the time required persample). 3 j j The diluted sample is next pumped by pump 50 through injector 51 into the measuring flowcell 52. The sample stream is confined by a fluid sheath provided by pump 54 of liquid from supply 56in order to obtain a narrow, rapidly flowing sample stream. 1 As noted, flow cell 52 is constructed so that fluid is introduced in one stream through'injectornozzle 51 and in an annular stream,surrounding the first stream, by pump '54 into annular space 53. The velocities of the central sample stream and the annular or sheath stream are controlled such that laminar flow conditions are established at the junction of the two streams, hencethe two streams will move-together with the sheath stream effectively constricting the sample stream. The sheath fluid provided from source of supply 56 preferably is selected to provide the requisite viscosity which will permit laminar flow underthe head pressure provided by pump 54. It should also be selected so that there is close matching of refractive indices between the sheath and sample fluidls. 1 3 V I The'diluent from supply 48 and the sheath, fluid 56 may be the same if desired although the requirements for the two need not be identical. For example, it is desirable to control carefully the refractive index and/or vinyl, pyrrolidone and the like.0bviously the salt additive is usually and preferably simply NaCl.

The flow cell preferably has circular cross section and has the largest diameter adjacent the tip of injector nozzle 51, being tapered down stream from that point. To obtain a desirable center sample stream of microns in diameter, typically, the flow cell will be tapered down to an internal diameter of about 200 microns. By using such a flow cell the blood cells are transmitted along the central stream in single file, at high speeds. I I

The flow cell thus described then essentially confines the blood cells to a narrow stream wherein the blood cells move each through a particular point substantially one at a time and therefore each can be examined in sequence. Further because the center-stream confines the blood cells'to a substantially axial-flow, the latter motion of the blood cells is sharply limited and hence the cells will remain well within focus of an optical system. Thus the system of the invention includes an electro-optical subsystem which is shown'schemati'cally in FIG. 1. The subsystem preferably includes one or more optical devices 60 such'as lenses, mirrors and the like for illuminating separate portions of flow cell 52 with radiation from one or more sources; shown generically as spectral source 62. Typically associated with each optical device 60 is a detection device 64 for converting selected parameters of radiation from a corresponding device 60 into an electrical signal.

Forsimplicity in exposition only one detection device 64 is shown in FIG. 1 as electrically connected to other equipment. A typical detection device as shown in FIG. 2 includes radiation source 62A andan optical system, essentially an inverted microscope having eyepiece lens 66 disposed adjacent source 62A and objective lens 68 disposed adjacent one side of flow cell 52. Positioned intermediate lenses 66 and 68 is a mask 70 having a plurality of apertures such as slots 72 therein. The microscope formed of lenses 66 and 68 is so positioned that radiation from source 62A, formed into a plurality V of discrete apparent sources by lens 66 and mask 70 is viscosity of both of the fluids, as well as the osmotic pressure produced across the surface of sample cells which may be suspended in or associated with the fluids. To these ends, the fluids are preferably aqueous solutionsjilntaining both additives which are polymeric and ad i lves which'are salts. The control of refractive index is established by adjusting the concentration of the polymer in the solution. For a given concentration of polymer, the viscosity of the fluid can be adjusted by selectirlg an appropriate degree of polymerization (i.e. the average molecular weight of the polymer) which parameter has relatively little effect on refractive index. Lastly, the polymer will have little effect on osmotic pressure, so the fluid may include a complementary dissolved salt, the concentration of which will serve to adjust the osmotic pressure to some desired value. v

Thus typical polymeric additives for ent and sheath fluids are polyethylene glycol and'the like, and blood plasma extenders such as dextran, poly use in the dilu-.

focused by objective lens 68 at a like plurality of spots distributed axially substantially along the center of the center stream in flow cell 52.

Disposed on the opposite side of flow cell 52 and typically on the common optic axis of lenses 66 and 68 ia another lens 74, typically a microscope type objective which has an equal orhigher numerical aperture than lens 68, and therefore is capable of accepting all of the radiation transmitted through lens 68 from source 62A.

Disposed between lens 74 and its focal planeis apertured diaphragm 76. Apertu'res 78 in the latter are disposed such that light originating from each aperture 72 in mask is substantially focused through a corresponding aperture 78 in diaphragm 76. Disposed on the opposite side of each aperture 78 from lens 74 are optiparticular wavelengths as desired. Positioned to detect the radiation transmitted by each of filters 80 are corresponding ones of detectors 82. The lattertypically can 7 be photodiodes with extremely fast rise times, or photocells of other known types.

In the optical system as described, typically the source can be means for providing a specified spectrum, such as a xenon high intensity lamp, with or without a selected output filter. It is quite important to set up a number of images (corresponding to the number of apertures 72 in mask 70) and that these images be rather close to one another axially along the center of flow cell 52. This structure serves to minimize the time for a single cell to go from one image or lightspot to the next and to activate corresponding detectors 82, thereby serving to minimize consequences of the errors in the velocity of the cells traversing flow cell 52. This minimization of velocity error is particularly important in instances where it is desired to correlate successive readings of the same cell so as to characterize the cell according to several different measurements.

It will also be appreciated that the structure described in FIG. 2 is particularly useful in determining either absorption or transmission characteristics of a blood cell. To measure scattering, a typical system is shown in FIG. 3 as including source 623, apertured mask 71 and lens 68. These elements are positioned to focus the radiation from the aperture in mask 71 as a spot centered within flow cell 52. The system of FIG. 3 also includes lens 74, screen 75 and detector 82. Screen 75 is simply an opaque screen with an annular opening 77 therein. Opening 77 of course surrounds an opaque center. As well known in the art, lens 74, screen 75 and detector 82 are so positioned with respect to one another and to flow cell 52, that light scattered from a particle in flow cell 52 over some given range of angles will be detected by detector 82.

Referring again to FIG. 1, it will be seen that typically, there are at least two output lines 84 and 85 from respective detectors 82 in a detection device 64. The signal transmitted on line 84 is designated as X. Similarly the signal being transmitted on line 85 is designated as Y. It can be assumed that the signals X and Y each represent a different shape-dependent function of a particular cell occasioned by an appropriate selection of input radiation to the cell, the selected output radiation from the cell, the position of the input radiation to the cell and the position of the detector and detector optics with respect to the output radiation from the cell. Line 84 is coupled to provide the signal X to the input of function element 86 which is capable of generating from signal X an output signal which is an exponential in the form X"" where n and m are values selected according to the type of shape-dependent functions that X and Y may respectively be. Function element 86 may be any of a number of known types of electronic elements. For example, such an element may be a diode function generator the output current of which is an arbitrary function of an input voltage. Such diode function generators are commercially available as Model SPFX-N/P circuits currently sold by Teledyne Philbrick/Nexus, Dedham, Massachusetts and described in Teledyne Philbrick/Nexus Bulletin EEM, File No. 1100.

The output of function element 86 is connected to one input of a ratiometric device 88, another input of device 88 being connected to line 85. Obviously, if line 85 provides a voltage and the output of element 86 is a current, appropriate voltage/current conversion equipment should be introduced in at least one of the lines so that the two inputs to ratiometric device 88 are similar parameters. Ratiometric device 88 is a wellknown device capable of accepting a pair of different inputs and for providing an output which is a ratio of those inputs. The output of ratiometer 88 is typically in a formflX)/f(Y) where X is X""". It will be appreciated that the output of the ratiometer is therefore some shape factor for the particle. Obviously, and strictly speaking, the two inputs to ratiometer 88 will often not be simultaneous if the corresponding detectors are triggered in sequence. Hence, the two inputs should be time correlated, as by introducing a delay line into line 84 or by employing known sample-and-hold techniques. Such refinements have been omitted from the drawings for the sake of simplicity in exposition and because such expedients are so well known in the art.

In order to employ the shapefactor or output signal from ratiometric device 88 for classifying blood cells, a typical circuit is shown in FIG. 1 wherein the output from ratiometric device 88 is connected in parallel to respective first inputs of a group of comparators 90, 92 and 94. Each such comparator includes a second input which is connected to a respective corresponding source, 96, 97 and 98 of reference signals. The output of comparators 90, 92 and 94 are respectively connected to counters 100, 101 and 102. Comparators 90, 92 and 94 are of the type known as threshholding comparators for providing an output pulse only when the input signal is greater than (or less than as the case may be) the amplitude of the reference signal provided from a corresponding reference source. Such comparators are also well known in the art and need not further be described here.

In operation of the electro-optical system as heretofore described, the shape factor y is derived as the ratio of any two of a number of different shape dependent functions. I-Ience, as a blood cell traverses flow cell 52, at least two shape dependent functions of that cell are then measured.

For example, one can select as the two desired shape dependent functions, the mean effective thickness of the nucleus of the cell and the nuclear volume. To determine the value of these functions, one then need simply measure light absorption by nucleic acid in two selected wavelength regions. For one of these regions (ca. 2 m p. for DNA) the nuclei of white blood cells are usually optically dense, while for some other wavelengths, the nuclei are optically thin. These wavelength regions can be selected by corresponding choices of filters 80and the spectral output of source 62A of FIG. 2. Self-shadowing will therefore make the ratio of these two measurements a non-linear, thickness-dependent value. This ratio of measurement provides a measure of optical depth and therefore of the mean effective thickness of the nucleus. In such case, the circuit of FIG. 1 should be modified as shown in FIG. 4 to provide ratiometric circuit 104 connected to lines 84 and for determining the ratio of X/Y. The output of ratiometric circuit 104 is then connected as the input to function element 86. The output of element 86 can be considered as flZ) or Z"""; the output of ratiometric circuit can be considered as the term flX)/f(Y). Hence the output of ratiometric device 88 will be f (Z')/f(Y) where Z Z"''. The measurement done in the wavelength region where the nucleus is optically thin provides the signal Y which is proportional to the nuclear volume. This latter measurement is correlated in device 9 88 ashe'retofore described with the thickness data from function element 86 to give a shape describing factor.

' Measurements for an optically thick absorber are usually orientation dependent, i.e. even if the radiation incident on the cell is of constant amplitude, theextent of absorption noted by the detector depends markedly on the spatial orientation of the nuclear components unless for example the latter are grouped symmetrically as a sphere. I

In order to minimize orientation effects, the cell should be illuminated by a system from a number of different directions (referred to the cell itself or as seen by the cell) the system being arranged so that the sum of the signals corresponding to the observed cellular or nuclear cross-section is substantially independent of orientation, i.e. is substantially invariant. Ideally there are at least three such directions, ideally mutually orthogonal and at least 60 mutually separated from one another. Thus, for example, in the simplest form as shown generally in FIG. three individual sources 162A, 1628 and 162C are provided and disposed to diy rect three corresponding light beams along mutually orthogonal paths to intersect at a common point 163 preferably disposed on the longitudinal axis of flowcell 52. Similarly, three detection devices 164A, 1648 and 164C are disposed to detect radiation emitted by, for example, the nucleus of a cell located at point 163 responsively respectively toirradiation bymeans from sources 162A, 162B and 162C. The detection devices, being adapted to provide electrical output signals responsively to radiation incident thereon, all have their output terminals connected to summing device 166. The signal fromdevice 166, the sum of all the inputs, will be substantially invariant for a particular aggregate of particles of given shape and size located at point 163 regardless of the orientation of the particle of aggregate.

It will be apparent that the beams from sources 162A, 1623 and 162C need not intersect but can illuminate successively a blood cell traversing flowcell52 assuming that the orientation of that blood cell with respect to the sources does not materially change. In such case, the output signals from detection devices 164A,- ll64B and l6dC can be suitably delayed or stored so as to be summed later.

Alternatively, one can provide means, such as a spiralled input channel. to flowcell 52, for introducing a known rotation about one or more axes to blood cells traveling down the axis of the flowcell. In such case, if only one light source is used and is focussed to a sufficiently large spot the blood cell can at successive intervals see the light beam from at least two, and preferably three, mutually orthogonal directions. In such case a simple detector, responsive to correspondingly successive levels of radiation outputs from the blood cell will suffice to feed a device which ultimately will sum the successive signals.

It will be apparent that the systems just described eliminate variations due to random orientation only in the case of objects of simple geometry such as ellipsoids, cylinders, and the like. The less symmetric an object is, the more directions of illumination and of observation will be required to eliminate orientation effects substantially. In the limit, hemispheric illumination of the object, together with hemispheric collection of the radiation output, completely eliminates variations of the output due to orientation for even the most complex shape. A system intended to approach this ideal performance is shown in FIG. 6 wherein the illumination provided by lens 68 should be at a very high aperture (e.g. ideally an illuminating cone approaching a 180 aperture. This can be achieved as shown by using immersion techniques whereby the interspace between the objective lens and the wall of the flowcell is filled with fluid 69 having a high index of refraction (e.g., 1), preferably matched to the index of refraction of the carrier fluid in flowcell 52. Cones in excess of can be formedin this manner. For a wide angle illuminating cone of uniform radiance, the interaction between an absorber of arbitrary shape and the incoming radiation approaches orientation invariancy. This occurs because regardless of the absorbers position, the amount of light coming at it from a given relative angle of approach is always the same. Should the absorber rotate, a given angle of incidence contains different rays,-'but the total intensity of the bundle of rays should be constant.

A shape factor can also be obtained by instead measuring the nuclear volume and the nuclear surface and correlating these measurements to obtain a shape factor. The nuclear volume, asnoted, is best determined photometrically throughmeasurements of its DNA content. In addition to measuring by direct absorption of radiation in a region where the nucleus is optically thin, the nuclear volume may also be determined by measuring the fluorescent reemissionof the absorbed light. This may be preferablev in some cases because to obtain good signal-to-noise ratio in absorption measurements a substantial fraction of the light must be absorbed. For fluroescent measurements, on the other hand, the same signal-to-noise ratio can be obtained with much smaller signals. Given a good conversion efficiency between absorption ans subsequent fluorescent ernission, substantially less absorption is needed in fluorescence measurements. Low absorption is desirable since it reduces the effect of self-shadowing of one portion of the nucleus by another, which tends to distort the relationship between measurements and the total DNA content. Fluorescent measurements also tend to be less orientation sensitive with respect to nonspherical nuclei.

The surface related magnitude can be measured in terms of light scattering from the molecular structure of the nucleus where the particles are in the order of a few wavelengths in size. Because the extent of scattering produced per unit particle volume increases as the particle becomes smaller an assemblage of small particles will scatter more light than a single particle of the same body. However, in such case spurious scattering must be eliminated; Such spurious scattering usually arises from three sources: (1 scattering from the blood cell cytoplasm which may be eliminated by matching the index of refraction of the fluid carrying the blood cell to the cytoplasm by the techniques heretofore described; (2) scattering from the blood cell membrane which is particularly strong for fairly thick membranes such as are found in erythrocytes; and (3) colloid scattering from the whole cell and surrounding solvent. Scattering from erythrocytes can be eliminated by providing a detector 64 which will detect hemoglobin, and by using the output of such detector to gate out any signals arising from erythrocytes. Colloid scattering is very strongly concentrated in the forward direction and thus the provision of detector optics of the type to avoid detecting wholly forward scattered light will serve to minimize the effect of colloid scattering.

It is appreciated that the scattering behavior of small particles of arbitrary shape and size in the order of a few wavelengths is exceedingly complex. However, the shape dependent factor sought has no particular linear or accuracy requirement, so that as long as scattering efficiency grows for smaller particles, one can differentiate between a large particle of given volume and an assemblage of smaller particles having the same total volume. Therefore, the only requirement for the sizedependent function based on scattering is that it have a slope which does not change in sign. Particularly when using scattering techniques for establishing a shape dependent factor, it is important to employ a broadband source, ie a source which provides radiation covering at least one octave of wavelengths at amplitudes above some minimum level. Of course the detector employed should be also responsive substantially across the bandwidth of the source. By employing such a broadband source and broadband detector one achieves smoothing of the scattering function'preferably a curve with substantially invariant sign of its slope.

Once the two functions, X and Y, (or Z and Y as the case may be) are obtained, normalized by function element 86 and ratioed by 88, a nuclear shape factor can be determined for each white blood cell traversing flow cell 52. By comparing the magnitude'(in comparators 90, 92 and 94) of each such factor against a corresponding different reference magnitude from a corresponding reference source, one then can arbitrarily count those shape factors which for example represent very spherical, very unspherical and medium sphericity" as an example of an arbitrary classification scheme. Such classification is of course not limited necessarily to any number of classes. Indeed, the outputs of the comparators need not be directly counted but can be gated to counters according to correlation with yet other aspects of the blood cells as may be determined for example by yet another detector devices 64.

Since certain changes may be made in the foregoing method and apparatus withough departing from the scope of the invention therein involved, it is intended that all matter contained in the above description and shown in the accompanying drawings be construed in an illustrative and not in a limiting sense.

What is claimed is:

1. In apparatus for determining parameters of a blood cell suspended in an axially moving flow stream, the improvement comprising in combination:

a first optical system for illuminating said cell with radiation;

a second optical system, having a numerical aperture equal to or greater than said first system, for gathering radiation emitted from said cell due to radiation incident thereon from said firstoptical system; and

detection means for producing a signal corresponding to the total radiation gathered by said second optical system, said first optical system being arranged for providing illumination to said cell from a plurality of different directions as seen by said cell so that said signal is substantially independent of the orientation of said cell.

2. The improvement as defined in claim 1 wherein said first optical system is formed to provide at least three independent beams of said radiation disposed substantially mutually orthogonally.

3. The improvement as defined in claim 2 wherein said second optical system comprises a plurality of optical light gathering devices corresponding in number to said light beams and each disposed to gather radiation from said cell due to the incidence on the latter of a respective one of said beams.

4. The improvement as defined in claim 3 including means for summing the total radiation gathered by said light gathering devices, and wherein said detection means comprises a detector substantially sensitive to said total radiation.

5. The improvement as defined in claim 3 wherein said detection means comprises a like plurality of detectors each being substantially sensitive to radiation gathered by a respective one of said light gathering devices so as to provide a corresponding signal, said improvement including means for summing all signals from said detectors.

6. The improvement as defined in claim 1 wherein said first optical system comprises means for directing radiation of a predetermined wavelength band in at least one cone of greater than to at least one focal spot disposed substantially on the axis of said flowstream,

said detection means being substantially sensitive to radiation across said band for providing a signal corresponding to detected radiation.

7. The improvement as defined in claim 6 wherein said means for directing radiation comprises a single refractor for directing a plurality of said cones of radiation to a corresponding plurality of focal spots disposed adjacent but separated from one another along the axis of said flowcell,

said detection means comprises a plurality of corresponding detectors substantially sensitive to radiation arising at respective ones of said focal spots, and

said second optical system comprises means for gathering radiation from each said focal spot and directing the gathered radiation from each spot to a respective one of said detectors. 

1. In apparatus for determining parameters of a blood cell suspended in an axially moving flow stream, the improvement comprising in combination: a first optical system for illuminating said cell with radiation; a second optical system, having a numerical aperture equal to or greater than said first system, for gathering radiation emitted from said cell due to radiation incident thereon from said first optical system; and detection means for producing a signal corresponding to the total radiation gathered by said second optical system, said first optical system being arranged for providing illumination to said cell from a plurality of different directions as seen by said cell so that said signal is substantially independent of the orientation of said cell.
 2. The improvement as defined in claim 1 wherein said first optical system is formed to provide at least three independent beams of said radiation disposed substantially mutually orthogonally.
 3. The improvement as defined in claim 2 wherein said second optical system comprises a plurality of optical light gathering devices corresponding in number to said light beams and each disposed to gather radiation from said cell due to the incidence on the latter of a respective one of said beams.
 4. The improvement as defined in claim 3 including means for summing the total radiation gathered by saId light gathering devices, and wherein said detection means comprises a detector substantially sensitive to said total radiation.
 5. The improvement as defined in claim 3 wherein said detection means comprises a like plurality of detectors each being substantially sensitive to radiation gathered by a respective one of said light gathering devices so as to provide a corresponding signal, said improvement including means for summing all signals from said detectors.
 6. The improvement as defined in claim 1 wherein said first optical system comprises means for directing radiation of a predetermined wavelength band in at least one cone of greater than 90* to at least one focal spot disposed substantially on the axis of said flowstream, said detection means being substantially sensitive to radiation across said band for providing a signal corresponding to detected radiation.
 7. The improvement as defined in claim 6 wherein said means for directing radiation comprises a single refractor for directing a plurality of said cones of radiation to a corresponding plurality of focal spots disposed adjacent but separated from one another along the axis of said flowcell, said detection means comprises a plurality of corresponding detectors substantially sensitive to radiation arising at respective ones of said focal spots, and said second optical system comprises means for gathering radiation from each said focal spot and directing the gathered radiation from each spot to a respective one of said detectors. 